U.S. patent number 7,471,972 [Application Number 11/021,162] was granted by the patent office on 2008-12-30 for sensor head for use with implantable devices.
This patent grant is currently assigned to DexCom, Inc.. Invention is credited to James H. Brauker, Rathbun Rhodes, Mark C. Shults, Mark A. Tapsak.
United States Patent |
7,471,972 |
Rhodes , et al. |
December 30, 2008 |
**Please see images for:
( Certificate of Correction ) ** |
Sensor head for use with implantable devices
Abstract
The present invention provides a sensor head for use in an
implantable device that measures the concentration of an analyte in
a biological fluid which includes: a non-conductive body; a working
electrode, a reference electrode and a counter electrode, wherein
the electrodes pass through the non-conductive body forming an
electrochemically reactive surface at one location on the body and
forming an electronic connection at another location on the body,
further wherein the electrochemically reactive surface of the
counter electrode is greater than the surface area of the working
electrode; and a multi-region membrane affixed to the nonconductive
body and covering the working electrode, reference electrode and
counter electrode. In addition, the present invention provides an
implantable device including at least one of the sensor heads of
the invention and methods of monitoring glucose levels in a host
utilizing the implantable device of the invention.
Inventors: |
Rhodes; Rathbun (Madison,
WI), Tapsak; Mark A. (San Diego, CA), Brauker; James
H. (San Diego, CA), Shults; Mark C. (Madison, WI) |
Assignee: |
DexCom, Inc. (San Diego,
CA)
|
Family
ID: |
25437709 |
Appl.
No.: |
11/021,162 |
Filed: |
December 22, 2004 |
Prior Publication Data
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Document
Identifier |
Publication Date |
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US 20050103625 A1 |
May 19, 2005 |
|
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
Issue Date |
|
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09916711 |
Jul 27, 2001 |
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Current U.S.
Class: |
600/347;
600/365 |
Current CPC
Class: |
A61B
5/14532 (20130101); A61B 5/14865 (20130101); C12Q
1/002 (20130101); G01N 33/48785 (20130101); G01N
27/3272 (20130101); G01N 33/48707 (20130101); C12Q
1/006 (20130101) |
Current International
Class: |
A61B
5/05 (20060101) |
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|
Primary Examiner: Nasser; Robert L
Attorney, Agent or Firm: Knobbe Martens Olson & Bear
LLP
Parent Case Text
This application is a continuation of Ser. No. 09/916,711, filed
Jul. 27, 2001 now abandoned.
Claims
What is claimed is:
1. A sensor for use in a glucose measuring device, the sensor
comprising: a first electrode, a second electrode, and a
non-conductive body located between the first electrode and the
second electrode, wherein the first electrode and the second
electrode each form an electrochemically reactive surface at one
end of the sensor and an electronic connection at another end of
the sensor, wherein an electrochemically reactive surface area of
the second electrode is greater than an electrochemically reactive
surface area of the first electrode; and a multi-region membrane
covering the first electrode and the second electrode, wherein the
multi-region membrane comprises an immobilized enzyme domain
comprising an enzyme in at least a portion thereof, and wherein the
immobilized enzyme is deactivated over the electrochemical reactive
surface area of second electrode.
2. The sensor of claim 1, wherein a portion of the multi-region
membrane covering the second electrode further comprises an
additional domain capable of selectively blocking glucose from an
area over the electrochemically reactive surface of the second
electrode.
3. The sensor of claim 1, wherein the multi-region membrane
comprises an oxygen antenna domain.
4. The sensor of claim 1, wherein the multi-region membrane
comprises a first region distant from the electrochemically
reactive surfaces, a second region less distant from the
electrochemically reactive surfaces than the first region and a
third region less distant from the electrochemically reactive
surfaces than the second region.
5. The sensor of claim 4, wherein the first region comprises a cell
impermeable domain.
6. The sensor of claim 4, wherein the second region includes a
glucose exclusion domain that does not substantially cover the
electrochemically reactive surface of the first electrode.
7. The sensor of claim 4, wherein the third region comprises the
immobilized enzyme domain distant from the electrochemically
reactive surfaces and an electrolyte-containing domain less distant
from the electrochemically reactive surfaces.
8. The sensor of claim 4, wherein the third region comprises the
immobilized enzyme domain distant from the electrochemically
reactive surfaces and a resistance domain more distant from the
electrochemically reactive surfaces than the immobilized enzyme
domain.
9. The sensor of claim 4, wherein the first region is permeable to
oxygen and glucose.
10. The sensor of claim 4, wherein the second region is permeable
to oxygen and interferes with glucose transport across the
membrane, further wherein the second region does not cover the
electroehemically reactive surface of the first electrode.
11. The sensor of claim 1, wherein the first electrode comprises a
working electrode.
12. The sensor of claim 11, wherein the second electrode comprises
a reference electrode or counter electrode.
13. A sensor for use in a glucose measuring device, the sensor
comprising: a first electrode, a second electrode, and a
non-conductive body located between the first electrode and the
second electrode, wherein the first electrode and the second
electrode each form an electrochemically reactive surface at one
end of the sensor and an electronic connection at another end of
the sensor; and a multi-region membrane covering the first
electrode and the second electrode, wherein the multi-region
membrane comprises an immobilized enzyme domain comprising an
active enzyme in at least a portion thereof, and wherein the
electrochemical reactive surface area of the second electrode is
situated at a location distant from the active enzyme. wherein the
immobilized enzyme is deactivated over the second electrode.
14. The sensor of claim 13, wherein the multi-region membrane
comprises a first region distant from the electrochemically
reactive surfaces and a further region less distant from the
electrochemically reactive surfaces.
15. The sensor of claim 14, wherein the first region comprises a
cell impermeable domain.
16. The sensor of claim 14, wherein the further region comprises
the immobilized enzyme domain distant from the electrochemically
reactive surfaces and an electrolyte-containing domain less distant
from the electrochemically reactive surfaces than the enzyme
domain.
17. The sensor of claim 16, wherein the electrolyte-containing
domain comprises a hydrogel.
18. The sensor of claim 14, wherein the further region further
comprises a resistance domain more distant from the
electrochemically reactive surfaces than the immobilized enzyme
domain.
19. The sensor of claim 14, wherein the further region comprises a
portion positioned over the second electrode and not positioned
over the first electrode, wherein the portion reduces consumption
of oxygen above the second electrode.
20. The sensor of claim 13, wherein the non-conductive body
comprises at least one of a plastic or a polymer.
21. The sensor of claim 13, wherein the first electrode comprises a
working electrode.
22. The sensor of claim 21, wherein the second electrode comprises
a reference electrode or counter electrode.
Description
FIELD OF THE INVENTION
The present invention relates generally to novel sensor heads
utilized with implantable devices, devices including these sensor
heads and methods for determining analyte levels using these
implantable devices. More particularly, the invention relates to
sensor heads, implantable devices including these sensor heads and
methods for monitoring glucose levels in a biological fluid using
these devices.
BACKGROUND OF THE INVENTION
Amperometric electrochemical sensors require a counter electrode to
balance the current generated by the species being measured at the
working electrode. In the case of a glucose oxidase based glucose
sensor, the species being measured at the working electrode is
H.sub.2O.sub.2. Glucose oxidase catalyzes the conversion of oxygen
and glucose to hydrogen peroxide and gluconate according to the
following reaction: Glucose+O.sub.2
.fwdarw.Gluconate+H.sub.2O.sub.2
Because for each glucose molecule metabolized, there is a
proportional change in the product H.sub.2O.sub.2, one can monitor
the change in H.sub.2O.sub.2 to determine glucose concentration.
Oxidation of H.sub.2O.sub.2 by the working electrode is balanced by
reduction of ambient oxygen, enzyme generated H.sub.2O.sub.2, or
other reducible species at the counter electrode. In vivo glucose
concentration may vary from about one hundred times or more that of
the oxygen concentration. Consequently, oxygen becomes a limiting
reactant in the electrochemical reaction and when insufficient
oxygen is provided to the sensor, the sensor will be unable to
accurately measure glucose concentration. Those skilled in the art
have come to interpret oxygen limitations resulting in depressed
function as being a problem of availability of oxygen to the
enzyme.
As shown in FIG. 1, the sensor head 10 includes a working electrode
21 (anode), counter electrode 22 (cathode), and reference electrode
20 which are affixed to the head by both brazing 26 the electrode
metal to the ceramic and potting with epoxy 28. The working
electrode 21 (anode) and counter-electrode 22 (cathode) of a
glucose oxidase-based glucose sensor head 10 require oxygen in
different capacities. Prior art teaches an enzyme-containing
membrane that resides above an amperometric electrochemical sensor.
In FIG. 1, region 32 includes an immobilized enzyme, i.e. glucose
oxidase. Within the enzyme layer above the working electrode 21,
oxygen is required for the production of H.sub.2O.sub.2 from
glucose. The H.sub.2O.sub.2 produced from the glucose oxidase
reaction further reacts at surface 21a of working electrode 21 and
produces two electrons. The products of this reaction are two
protons (2H.sup.+), two electrons (2e.sup.-), and one oxygen
molecule (O.sub.2) (Fraser, D. M. "An Introduction to In Vivo
Biosensing: Progress and problems." In "Biosensors and the Body,"
D. M. Fraser, ed., 1997, pp. 1-56 John Wiley and Sons, New York).
In theory, the oxygen concentration near the working electrode 21,
which is consumed during the glucose oxidase reaction, is
replenished by the second reaction at the working electrode.
Therefore, the net consumption of oxygen is zero. In practice,
neither all of the H.sub.2O.sub.2 produced by the enzyme diffuses
to the working electrode surface nor does all of the oxygen
produced at the electrode diffuse to the enzyme domain.
With further reference to FIG. 1, the counter electrode 22 utilizes
oxygen as an electron acceptor. The most likely reducible species
for this system are oxygen or enzyme generated peroxide (Fraser, D.
M. supra). There are two main pathways by which oxygen may be
consumed at the counter electrode 22. These are a four-electron
pathway to produce hydroxide and a two-electron pathway to produce
hydrogen peroxide. The two-electron pathway is shown in FIG. 1.
Oxygen is further consumed above the counter electrode by the
glucose oxidase in region 32. Due to the oxygen consumption by both
the enzyme and the counter electrode, there is a net consumption of
oxygen at the surface 22a of the counter electrode. Theoretically,
in the domain of the working electrode there is significantly less
net loss of oxygen than in the region of the counter electrode. In
addition, there is a close correlation between the ability of the
counter electrode to maintain current balance and sensor function.
Taken together, it appears that counter electrode function becomes
limited before the enzyme reaction becomes limited when oxygen
concentration is lowered.
Those practicing in the field of implantable glucose oxidase
sensors have focused on improving sensor function by increasing the
local concentration of oxygen in the region of the working
electrode. (Fraser, D. M. supra).
We have observed that in some cases, loss of glucose oxidase sensor
function may not be due to a limitation of oxygen in the enzyme
layer near the working electrode, but may instead be due to a
limitation of oxygen at the counter electrode. In the presence of
increasing glucose concentrations, a higher peroxide concentration
results, thereby increasing the current at the working electrode.
When this occurs, the counter electrode limitation begins to
manifest itself as this electrode moves to increasingly negative
voltages in the search for reducible species. When a sufficient
supply of reducible species, such as oxygen, are not available, the
counter electrode voltage reaches a circuitry limit of -0.6V
resulting in compromised sensor function (see FIG. 3).
FIG. 3 shows simultaneous measurement of counter-electrode voltage
and sensor output to glucose levels from a glucose sensor implanted
subcutaneously in a canine host. It can be observed that as glucose
levels increase, the counter electrode voltage decreases. When the
counter electrode voltage reaches -0.6V, the signal to noise ratio
increases significantly. This reduces the accuracy of the device.
FIG. 4 shows a further example of another glucose sensor in which
the counter-electrode reaches the circuitry limit. Again, once the
counter electrode reaches -0.6V, the sensitivity and/or signal to
noise ratio of the device is compromised. In both of these
examples, glucose levels reached nearly 300 mg/dl. However, in FIG.
3 the sensor showed a greater than three-fold higher current output
than the sensor in FIG. 4. These data suggest that there may be a
limitation of reducible species at the counter electrode, which may
limit the sensitivity of the device as the glucose levels increase.
In contrast, FIG. 5 shows a glucose sensor in which the counter
electrode voltage did not reach -0.6V. In FIG. 5 it can be observed
that the sensor was able to maintain a current balance between the
working and counter electrodes, thereby enabling accurate
measurements throughout the course of the experiment. The results
shown in FIGS. 3, 4 and 5 led the present inventors to postulate
that by keeping the counter electrode from reaching the circuitry
limit, one could maintain sensitivity and accuracy of the
device.
Two approaches have been utilized by others to relieve the counter
electrode limitation described above. The first approach involves
the widening of the potential range over which the counter
electrode can move in the negative direction to avoid reaching
circuitry limitations. Unfortunately, this approach increases
undesirable products that are produced at lower potentials. One
such product, hydrogen, may form at the counter electrode,which may
then diffuse back to the working electrode. This may contribute to
additional current resulting in erroneously high glucose
concentration readings. Additionally, at these increasingly
negative potentials, the probability of passivating or poisoning
the counter electrode greatly increases. This effectively reduces
the counter electrode surface area requiring a higher current
density at the remaining area to maintain current balance.
Furthermore, increased current load increases the negative
potentials eventually resulting in electrode failure.
The second approach is utilizing the metal case of the device as a
counter electrode (see U.S. Pat. No. 4,671,288, Gough or U.S. Pat.
No. 5,914,026, Blubaugh). This provides an initial excess in
surface area which is expected to serve the current balancing needs
of the device over its lifetime. However, when the counter
electrode reaction is a reduction reaction, as in Blubaugh, the
normally present metal oxide layer will be reduced to bare metal
over time leaving the surface subject to corrosion, poisoning, and
eventual cascade failure. This problem is magnified when
considering the various constituents of the body fluid that the
metal casing is exposed to during in vivo use. To date, there has
been no demonstration of long-term performance of such a device
with this counter electrode geometry.
Consequently, there is a need for a sensor that will provide
accurate analyte measurements, that reduces the potential for
cascade failure due to increasing negative potentials, corrosion
and poisoning, and that will function effectively and efficiently
in low oxygen concentration environments.
SUMMARY OF THE INVENTION
In one aspect of the present invention, a sensor head for use in a
device that measures the concentration of an analyte in a
biological fluid is provided that includes a non-conductive body; a
working electrode, a reference electrode and a counter electrode,
wherein the electrodes pass through the non-conductive body forming
an electrochemically reactive surface at one location on the body
and forming an electronic connection at another location on the
body, and further wherein the electrochemically reactive surface of
the counter electrode is greater than the surface area of the
working electrode; and a multi-region membrane affixed to the
nonconductive body and covering the working electrode, reference
electrode and counter electrode.
In another aspect of the present invention, a sensor head for use
in an implantable analyte measuring device is provided which
includes the same sensor head components as those described
above.
The sensor heads of the present invention include a multi-region
membrane that controls the number of species that are able to reach
the surface of the electrodes. In particular, such a membrane
allows the passage of desired substrate molecules (e.g. oxygen and
glucose) and rejects other larger molecules that may interfere with
accurate detection of an analyte. The sensor heads of the present
invention also provide a larger counter electrode reactive surface
that balances the current between the working and counter
electrodes, thereby minimizing negative potential extremes that may
interfere with accurate analyte detection.
In another aspect of the present invention, an implantable device
for measuring an analyte in a biological fluid is provided
including at least one of the sensor heads described above. In
still another aspect of the present invention, a method of
monitoring glucose levels is disclosed which includes the steps of
providing a host, and an implantable device as provided above and
implanting the device in the host.
Further encompassed by the invention is a method of measuring
glucose in a biological fluid including the steps of providing a
host and a implantable device described above, which includes a
sensor head capable of accurate continuous glucose sensing; and
implanting the device in the host.
The sensor head, membrane architectures, devices and methods of the
present invention allow for the collection of continuous
information regarding desired analyte levels (e.g. glucose). Such
continuous information enables the determination of trends in
glucose levels, which can be extremely important in the management
of diabetic patients.
Definitions
In order to facilitate an understanding of the present invention, a
number of terms are defined below.
The term "sensor head" refers to the region of a monitoring device
responsible for the detection of a particular analyte. The sensor
head generally comprises a non-conductive body, a working electrode
(anode), a reference electrode and a counter electrode (cathode)
passing through and secured within the body forming an
electrochemically reactive surface at one location on the body and
an electronic connective means at another location on the body, and
a multi-region membrane affixed to the body and covering the
electrochemically reactive surface. The counter electrode has a
greater electrochemically reactive surface area than the working
electrode. During general operation of the sensor a biological
sample (e.g., blood or interstitial fluid) or a portion thereof
contacts (directly or after passage through one or more membranes
or domains) an enzyme (e.g., glucose oxidase); the reaction of the
biological sample (or portion thereof) results in the formation of
reaction products that allow a determination of the analyte (e.g.
glucose) level in the biological sample. In preferred embodiments
of the present invention, the multi-region membrane further
comprises an enzyme domain, and an electrolyte phase (i.e., a
free-flowing liquid phase comprising an electrolyte-containing
fluid described further below).
The term "analyte" refers to a substance or chemical constituent in
a biological fluid (e.g., blood, interstitial fluid, cerebral
spinal fluid, lymph fluid or urine) that can be analyzed. A
preferred analyte for measurement by the sensor heads, devices and
methods of the present invention is glucose.
The term "electrochemically reactive surface" refers to the surface
of an electrode where an electrochemical reaction takes place. In
the case of the working electrode, the hydrogen peroxide produced
by the enzyme catalyzed reaction of the analyte being detected
reacts creating a measurable electronic current (e.g. detection of
glucose analyte utilizing glucose oxidase produces H.sub.2O.sub.2
peroxide as a by product, H.sub.2O.sub.2 reacts with the surface of
the working electrode producing two protons (2H.sup.+), two
electrons (2e.sup.-) and one molecule of oxygen (O.sub.2) which
produces the electronic current being detected). In the case of the
counter electrode, a reducible species, e.g. O.sub.2 is reduced at
the electrode surface in order to balance the current being
generated by the working electrode.
The term "electronic connection" refers to any electronic
connection known to those in the art that may be utilized to
interface the sensor head electrodes with the electronic circuitry
of a device such as mechanical (e.g., pin and socket) or
soldered.
The term "domain" refers to regions of the membrane of the present
invention that may be layers, uniform or non-uniform gradients
(e.g. anisotropic) or provided as portions of the membrane.
The term "multi-region membrane" refers to a permeable membrane
that may be comprised of two or more domains and constructed of
biomaterials of a few microns thickness or more which are permeable
to oxygen and may or may not be permeable to glucose. One of the
the membranes may be placed over the sensor body to keep host cells
(e.g., macrophages) from gaining proximity to, and thereby
damaging, the enzyme membrane or forming a barrier cell layer and
interfering with the transport of analyte across the tissue-device
interface.
The phrase "distant from" refers to the spatial relationship
between various elements in comparison to a particular point of
reference. For example, some embodiments of a biological fluid
measuring device comprise a multi-region membrane that may be
comprised of a number of domains. If the electrodes of the sensor
head are deemed to be the point of reference, and one of the
multi-region membrane domains is positioned farther from the
electrodes, than that domain is distant from the electrodes.
The term "oxygen antenna domain" and the like refers to a domain
composed of a material that has higher oxygen solubility than
aqueous media so that it concentrates oxygen from the biological
fluid surrounding the biointerface membrane. The domain can then
act as an oxygen reservoir during times of minimal oxygen need and
has the capacity to provide on demand a higher oxygen gradient to
facilitate oxygen transport across the membrane. This enhances
function in the enzyme reaction domain and at the counter electrode
surface when glucose conversion to hydrogen peroxide in the enzyme
domain consumes oxygen from the surrounding domains. Thus, this
ability of the oxygen antenna domain to apply a higher flux of
oxygen to critical domains when needed improves overall sensor
function.
The term "solid portions" and the like refer to a material having a
structure that may or may not have an open-cell configuration but
in either case prohibits whole cells from traveling through or
residing within the material.
The term "substantial number" refers to the number of cavities or
solid portions having a particular size within a domain in which
greater than 50 percent of all cavities or solid portions are of
the specified size, preferably greater than 75 percent and most
preferably greater than 90 percent of the cavities or solid
portions have the specified size.
The term "co-continuous" and the like refers to a solid portion
wherein an unbroken curved line in three dimensions exists between
any two points of the solid portion.
The term "host" refers to both humans and animals.
The term "accurately" means, for example, 90% of measured glucose
values are within the "A" and "B" region of a standard Clarke error
grid when the sensor measurements are compared to a standard
reference measurement. It is understood that like any analytical
device, calibration, calibration validation and recalibration are
required for the most accurate operation of the device.
The phrase "continuous glucose sensing" refers to the period in
which monitoring of plasma glucose concentration is continuously
performed, for example, about every 10 minutes.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 Illustration of thermodynamically favored reactions at the
working electrode and counter electrode at the desired voltage
potentials.
FIG. 2A depicts a cross-sectional exploded view of a sensor head of
the present invention wherein the multi-region membrane comprises
three regions.
FIG. 2B depicts a cross-sectional exploded view of a sensor head of
the present invention wherein a portion of the second membrane
region does not cover the working electrode.
FIG. 2C depicts a cross-sectional exploded view of a sensor head of
the present invention which includes two distinct regions, wherein
the region adjacent the electrochemically reactive surfaces
includes a portion positioned over the counter electrode which
corresponds to a silicone domain.
FIG. 2D depicts a cross-sectional exploded view of a sensor head of
the present invention wherein an active enzyme of the immobilized
enzyme domain is positioned only over the working electrode.
FIG. 2E depicts a cross-sectional exploded view of a sensor head of
the present invention wherein the enzyme positioned over the
counter electrode has been inactivated.
FIG. 2F depicts a cross-sectional exploded view of a sensor head of
the present invention wherein the membrane region containing
immobilized enzyme is positioned only over the working
electrode.
FIG. 3 Illustration of an implantable glucose sensor's ability to
measure glucose concentration during an infusion study in a canine
when the counter electrode voltage drops to the electronic
circuitry limit at approximately 0.75 hours wherein the sensor
current output reaches 2.50 nA.
FIG. 4 Illustration of an implantable glucose sensor's ability to
measure glucose concentration during an infusion study in a canine
when the counter electrode voltage drops to the electronic
circuitry limit after 0.5 hours wherein the sensor current output
reaches 0.50 nA.
FIG. 5 Illustration of an implantable glucose sensor's ability to
measure glucose concentration during an infusion study in a canine
when the counter electrode voltage is maintained above the
electronic circuitry limit.
FIG. 6A shows a schematic representation of a cylindrical analyte
measuring device including a sensor head according to the present
invention.
FIG. 6B is an exploded view of the sensor head of the device shown
in FIG. 6A.
FIG. 7 Graphical representation of the function of a device of the
present invention utilizing the multi-region membrane architecture
of FIG. 2B in vitro at 400 mg/dL glucose.
FIG. 8 depicts a cross-sectional exploded view of the electrode and
membrane regions of a prior sensor device where the electrochemical
reactive surface of the counter electrode is substantially equal to
the surface area of the working electrode.
FIG. 9 Graphical representation of the counter electrode voltage as
a function of oxygen concentration at 400 mg/dL glucose for sensor
devices including the membrane shown in FIG. 8.
DETAILED DESCRIPTION OF THE INVENTION
In a preferred embodiment, the sensor heads, devices and methods of
the present invention may be used to determine the level of glucose
or other analytes in a host. The level of glucose is a particularly
important measurement for individuals having diabetes in that
effective treatment depends on the accuracy of this
measurement.
The present invention increases the effectiveness of counter
electrode function by a method that does not depend on increasing
the local concentration of oxygen. In a preferred embodiment, the
counter electrode has an electrochemical reactive surface area
greater than twice the surface area of the working electrode
thereby substantially increasing the electrodes ability to utilize
oxygen as a substrate. Further enhancement of the counter
electrode's activity may be achieved if the electrode were made of
gold. In a second preferred embodiment, the counter electrode has a
textured surface, with surface topography that increases the
surface area of the electrode while the diameter of the electrode
remains constant. In a third preferred embodiment, the proximity of
the glucose oxidase enzyme to the counter electrode may be
decreased. Since the enzyme depletes oxygen locally, the counter
electrode would best be situated at a location distant from the
enzyme. This could be achieved by depleting the enzyme from or
inactivating the enzyme located in the region near and over the
counter electrode by methods known to those skilled in the art such
as laser ablation, or chemical ablation. Alternatively, the
membrane could be covered with an additional domain where glucose
is selectively blocked from the area over the counter
electrode.
In particular, the present invention reduces the potential for
electrode poisoning by positioning all electrodes underneath a
multi-region membrane so that there is control of the species
reaching the electrode surfaces. These membranes allow passage of
dissolved oxygen to support the counter electrode reactions at
reasonable negative potentials while rejecting larger molecules
which when reduced would coat the surface of the counter electrode
eventually leading to cascade failure. The positioning of the
counter electrode underneath the membrane assures that all currents
are passing through the same conductive media, thereby reducing
voltage losses due to membrane or solution resistance. In addition,
the counter electrode will be able to collect enough species for
the balancing current while minimizing the need to move towards
negative potential extremes.
Although the description that follows is primarily directed at
glucose monitoring sensor heads, devices and methods for their use,
the sensor heads, devices and methods of the present invention are
not limited to glucose measurement. Rather, the devices and methods
may be applied to detect and quantitate other analytes present in
biological fluids (including, but not limited to, amino acids and
lactate), especially those analytes that are substrates for oxidase
enzymes [see, e.g., U.S. Pat. No. 4,703,756 to Gough et al., hereby
incorporated by reference].
I. Nature of the Foreign Body Capsule
Devices and probes that are implanted into subcutaneous tissue will
almost always elicit a foreign body capsule (FBC) as part of the
body's response to the introduction of a foreign material.
Therefore, implantation of a glucose sensor results in an acute
inflammatory reaction followed by building of fibrotic tissue.
Ultimately, a mature FBC comprising primarily a vascular fibrous
tissue forms around the device (Shanker and Greisler, Inflammation
and Biomaterials in Greco R S, ed. Implantation Biology: The Host
Response and Biomedical Devices, pp 68-80, CRC Press (1994)).
In general, the formation of a FBC has precluded the collection of
reliable, continuous information, reportedly because of poor
vascularization (Updike, S. J. et al., "Principles of Long-term
Fully Implanted Sensors with Emphasis on Radiotelemetric Monitoring
of Blood Glucose from inside a Subcutaneous Foreign Body Capsule
(FBC)" in "Biosensors and the Body," D. M. Fraser, ed., 1997, pp.
117-38, John Wiley and Sons, New York). Thus, those skilled in the
art have previously attempted to minimize FBC formation by, for
example, using a short-lived needle geometry or sensor coatings to
minimize the foreign body.
In contrast to the prior art, the teachings of the present
invention recognize that FBC formation is the dominant event
surrounding long-term implantation of any sensor and must be
managed to support, rather than hinder or block, sensor
performance. It has been observed that during the early periods
following implantation of an analyte sensing device, particularly a
glucose sensing device, that glucose sensors function well.
However, after a few days to two or more weeks of implantation,
these device lose their function.
We have observed that this lack of sensor function is most likely
due to cells (barrier cells) that associate with the outer surface
of the device and physically block the transport of glucose into
the device (i.e. form a a barrier cell layer). Increased
vascularization would not be expected to overcome this blockage.
The present invention contemplates the use of particular
biointerface membrane architectures that interfere with barrier
cell layer formation on the membrane's surface. The present
invention also contemplates the use of these membranes with a
variety of implantable devices (e.g. analyte measuring devices
particularly glucose measuring devices).
II. The Sensor Head
In one embodiment of the sensor head of the invention, the body is
made of a non-conductive material such as ceramic, glass, or
polymer.
In a preferred embodiment, the sensor head interface region may
include several different layers and/or membranes that cover and
protect the electrodes of an implantable analyte-measuring device.
The characteristics of these layers and/or membranes are now
discussed in more detail. The layers and/or membranes prevent
direct contact of the biological fluid sample with the electrodes,
while permitting selected substances (e.g., analytes) of the fluid
to pass therethrough for reaction in an enzyme rich domain with
subsequent electrochemical reaction of formed products at the
electrodes.
It is well known in the art that electrode surfaces exposed to a
wide range of biological molecules may suffer poisoning of
catalytic activity and possible corrosion that could result in
failure. However, utilizing the unique multi-region membrane
architectures of the present invention, the active electrochemical
surfaces of the sensor electrodes are preserved, retaining activity
for extended periods of time in vivo. By limiting access to the
electrochemically reactive surface of the electrodes to a small
number of molecular species such as, for example, molecules having
a molecular weight of about 34 Daltons (the molecular weight of
peroxide) or less, only a small subset of the many molecular
species present in biological fluids are permitted to contact the
sensor. Use of such membranes has enabled sustained function of
devices for over one year in vivo.
A. Multi-Region Membrane
The multi-region membrane is constructed of two or more regions.
The multi-region membrane may be provided in a number of different
architectures. In one architecture, the multi-region membrane
includes a first region distant from the electrochemically reactive
surfaces, a second region less distant from the electrochemically
reactive surfaces and a third region adjacent to the
electrochemically reactive surfaces. The first region includes a
cell disruptive domain distant from the electrochemically reactive
surfaces and a cell impermeable domain less distant from the
electrochemically reactive surfaces. The second region is a glucose
exclusion domain and the third region includes a resistance domain
distant from the electrochemically reactive surfaces, an
immobilized enzyme domain less distant from the electrochemically
reactive surfaces, an interference domain less distant from the
electrochemically reactive surfaces than the immobilized enzyme
domain and a hydrogel domain adjacent to the electrochemically
reactive surfaces.
In another architecture, the multi-region membrane includes a first
region distant from the electrochemically reactive surfaces and a
further region less distant from the electrochemically reactive
surfaces. The first region includes a cell disruptive domain and a
cell impermeable domain as described above. The "further region"
includes a resistance domain, immobilized enzyme domain,
interference domain, and hydrogel domain and serves as the
equivalent of the "third region" described above. In certain
embodiments of the sensor head, the multi-region membrane further
includes an oxygen antenna domain. Each of these domains will now
be described in further detail.
i. Cell Disruptive Domain
The domain of the multi-region membrane positioned most distal to
the electrochemically reactive surfaces corresponds to the cell
disruptive domain. This domain includes a material that supports
tissue in-growth and may be vascularized. The cell disruptive
domain prevents formation of the barrier cell layer on the surface
of the membrane, which as described above, blocks the transport of
glucose into the sensor device. A useful cell disruptive domain is
described in a U.S. application entitled "Membrane for use with
Implantable Devices" which was filed on the same day as the present
application. The cell disruptive domain may be composed of an
open-cell configuration having cavities and solid portions. Cells
may enter into the cavities, however, they can not travel through
or wholly exist within the solid portions. The cavities allow most
substances to pass through, including, e.g., macrophages.
The open-cell configuration yields a co-continuous solid domain
that contains greater than one cavity in three dimensions
substantially throughout the entirety of the membrane. In addition,
the cavities and cavity interconnections may be formed in layers
having different cavity dimensions.
A linear line can be used to define a dimension across a cavity or
solid portion the length of which is the distance between two
points lying at the interface of the cavity and solid portion. In
this way, a substantial number of the cavities are not less than 20
microns in the shortest dimension and not more than 1000 microns in
the longest dimension. Preferably, a substantial number of the
cavities are not less than 25 microns in the shortest dimension and
not more than 500 microns in the longest dimension.
Furthermore, the solid portion has not less than 5 microns in a
substantial number of the shortest dimensions and not more than
2000 microns in a substantial number of the longest dimensions.
Preferably, the solid portion is not less than 10 microns in a
substantial number of the shortest dimensions and not more than
1000 microns in a substantial number of the longest dimensions and
most preferably, not less than 10 microns in a substantial number
of the shortest dimensions and not more than 400 microns in a
substantial number of the longest dimensions.
The solid portion may be made of polytetrafluoroethylene or
polyethylene-co-tetrafluoroethylene, for example. Preferably, the
solid portion includes polyurethanes or block copolymers and, most
preferably, includes silicone.
When non-woven fibers are utilized as the solid portion of the
present invention, the non-woven fibers may be greater than 5
microns in the shortest dimension. Preferably, the non-woven fibers
are about 10 microns in the shortest dimension and most preferably,
the non-woven fibers are greater than or equal to 10 microns in the
shortest dimension.
The non-woven fibers may be constructed of polypropylene (PP),
polyvinylchloride (PVC), polyvinylidene fluoride (PVDF),
polybutylene terephthalate (PBT), polymethylmethacrylate (PMMA),
polyether ether ketone (PEEK), polyurethanes, cellulosic polymers,
polysulfones, and block copolymers thereof including, for example,
di-block, tri-block, alternating, random and graft copolymers
(block copolymers are discussed in U.S. Pat. Nos. 4,803,243 and
4,686,044, hereby incorporated by reference). Preferably, the
non-woven fibers are comprised of polyolefins or polyester or
polycarbonates or polytetrafluoroethylene.
A subset of the cell disruptive domain is the oxygen antenna
domain. This domain can act as an oxygen reservoir during times of
minimal oxygen need and has the capacity to provide on demand a
higher oxygen gradient to facilitate oxygen transport across the
membrane. This domain may be composed of a material such as
silicone, that has higher oxygen solubility than aqueous media so
that it concentrates oxygen from the biological fluid surrounding
the biointerface membrane. This enhances function in the enzyme
reaction domain and at the counter electrode surface when glucose
conversion to hydrogen peroxide in the enzyme domain consumes
oxygen from the surrounding domains. Thus, this ability of the
oxygen antenna domain to apply a higher flux of oxygen to critical
domains when needed improves overall sensor function. Preferably,
this domain is composed of silicone and has a thickness of about
100 microns.
The thickness of the cell disruptive domain is usually not less
than about 20 microns and not more than about 2000 microns.
ii. Cell Impermeable Domain
The cell impermeable of the first region is positioned less distal
to the electrochemically reactive surfaces than the cell disruptive
domain of the same region. This domain is impermeable to host
cells, such as macrophages. Cell impermeable domains are described
in U.S. Pat. No. 6,001,067, herein incorporated by reference, and
in copending, commonly owned U.S. application entitled "Membrane
for use with Implantable Devices", Ser. No. 10/768,889, filed on
even date herewith. The inflammatory response that initiates and
sustains a FBC is associated with disadvantages in the practice of
sensing analytes. Inflammation is associated with invasion of
inflammatory response cells (e.g. macrophages) which have the
ability to overgrow at the interface and form barrier cell layers,
which may block transport of glucose across the biointerface
membrane. These inflammatory cells may also biodegrade many
artificial biomaterials (some of which were, until recently,
considered nonbiodegradable). When activated by a foreign body,
tissue macrophages degranulate, releasing from their cytoplasmic
myeloperoxidase system hypochlorite (bleach) and other oxidative
species. Hypochlorite and other oxidative species are known to
break down a variety of polymers, including ether based
polyurethanes, by a phenomenon referred to as environmental stress
cracking. Alternatively, polycarbonate based polyurethanes are
believed to be resistant to environmental stress cracking and have
been termed biodurable. In addition, because hypochlorite and other
oxidizing species are short-lived chemical species in vivo,
biodegradation will not occur if macrophages are kept a sufficient
distance from the enzyme active membrane.
The present invention contemplates the use of cell impermeable
biomaterials of a few microns thickness or more (i.e., a cell
impermeable domain) in most of its membrane architectures. This
domain of the biointerface membrane is permeable to oxygen and may
or may not be permeable to glucose and is constructed of biodurable
materials (e.g. for period of several years in vivo) that are
impermeable by host cells (e.g. macrophages) such as for example
polymer blends of polycarbonate based polyurethane and PVP.
The thickness of the cell impermeable domain is not less than about
10 microns and not more than about 100 microns.
iii. Glucose Exclusion Domain
The glucose exclusion domain includes a thin, hydrophobic membrane
that is non-swellable and blocks diffusion of glucose while being
permeable to oxygen. The glucose exclusion domain serves to allow
analytes and other substances that are to be measured or utilized
by the sensor to pass through, while preventing passage of other
substances. Preferably, the glucose exclusion domain is constructed
of a material such as, for example, silicone.
The glucose exclusion domain has a preferred thickness not less
than about 130 microns, more preferably not less than about 5 and
not more than about 75 microns and most preferably not less than 15
microns and not more than about 50 microns.
iv. Resistance Domain
In one embodiment of the sensor head the "third region" or "further
region" of the multi-region membrane includes a resistance domain.
When present, the resistance domain is located more distal to the
electrochemically reactive surfaces relative to other domains in
this region. As described in further detail below, the resistance
domain controls the flux of oxygen and glucose to the underlying
enzyme domain. There is a molar excess of glucose relative to the
amount of oxygen in samples of blood. Indeed, for every free oxygen
molecule in extracellular fluid, there are typically more than 100
glucose molecules present [Updike et al., Diabetes Care
5:207-21(1982)]. However, an immobilized enzyme-based sensor using
oxygen (O.sub.2) as cofactor must be supplied with oxygen in
non-rate-limiting excess in order to respond linearly to changes in
glucose concentration, while not responding to changes in oxygen
tension. More specifically, when a glucose-monitoring reaction is
oxygen-limited, linearity is not achieved above minimal
concentrations of glucose. Without a semipermeable membrane over
the enzyme domain, linear response to glucose levels can be
obtained only up to about 40 mg/dL; however, in a clinical setting,
linear response to glucose levels are desirable up to at least
about 500 mg/dL.
The resistance domain includes a semipermeable membrane that
controls the flux of oxygen and glucose to the underlying enzyme
domain (i.e., limits the flux of glucose), rendering the necessary
supply of oxygen in non-rate-limiting excess. As a result, the
upper limit of linearity of glucose measurement is extended to a
much higher value than that which could be achieved without the
resistance domain. The devices of the present invention contemplate
resistance domains including polymer membranes with
oxygen-to-glucose permeability ratios of approximately 200:1; as a
result, one-dimensional reactant diffusion is adequate to provide
excess oxygen at all reasonable glucose and oxygen concentrations
found in the subcutaneous matrix [Rhodes et al., Anal. Chem.,
66:1520-1529 (1994)].
In preferred embodiments, the resistance domain is constructed of a
polyurethane urea/polyurethane-block-polyethylene glycol blend and
has a thickness of not more than about 45 microns, more preferably
not less than about 15 microns, and not more than about 40 microns
and, most preferably, not less than about 20 microns, and not more
than about 35 microns.
v. Immobilized Enzyme Domain
When the resistance domain is combined with the cell-impermeable
domain, it is the immobilized enzyme domain which corresponds to
the outermost domain of the "third region" or "further region",
i.e. it is located more distal to the electrochemically reactive
surfaces as compared to the other domains in this region. In one
embodiment, the enzyme domain includes glucose oxidase. In addition
to glucose oxidase, the present invention contemplates the use of a
domain impregnated with other oxidases, e.g., galactose oxidase or
uricase, For an enzyme-based electrochemical glucose sensor to
perform well, the sensor's response must neither be limited by
enzyme activity nor cofactor concentration. Because enzymes,
including glucose oxidase, are subject to deactivation as a
function of ambient conditions, this behavior needs to be accounted
for in constructing sensors for long-term use.
Preferably, the domain is constructed of aqueous dispersions of
colloidal polyurethane polymers including the enzyme. Preferably,
the coating has a thickness of not less than about 2.5 microns and
not more than about 12.5 microns, preferably about 6.0 microns.
vi. Interference Domain
The interference domain in the "third region" or "further region"
is located less distant from the electrochemically reactive
surfaces than the immobilized enzyme domain in this same region. It
includes a thin membrane that can limit diffusion of molecular
weight species greater than 34 kD. The interference domain serves
to allow analytes and other substances that are to be measured by
the electrodes to pass through, while preventing passage of other
substances, including potentially interfering substances. The
interference domain is preferably constructed of a
polyurethane.
The interference domain has a preferred thickness of not more than
about 5 microns, more preferably not less than about 0.1 microns,
and not more than about 5 microns and, most preferably, not less
than about 0.5 microns, and not more than about 3 microns.
vii. Hydrogel Domain
The hydrogel domain is located adjacent to the electrochemically
reactive surfaces. To ensure electrochemical reaction, the hydrogel
domain includes a semipermeable coating that maintains
hydrophilicity at the electrode region of the sensor interface. The
hydrogel domain enhances the stability of the interference domain
of the present invention by protecting and supporting the membrane
that makes up the interference domain. Furthermore, the hydrogel
domain assists in stabilizing operation of the device by overcoming
electrode start-up problems and drifting problems caused by
inadequate electrolyte. The buffered electrolyte solution contained
in the hydrogel domain also protects against pH-mediated damage
that may result from the formation of a large pH gradient between
the hydrophobic interference domain and the electrode (or
electrodes) due to the electrochemical activity of the
electrode(s).
Preferably, the hydrogel domain includes a flexible,
water-swellable, substantially solid gel-like film having a "dry
film" thickness of not less than about 2.5 microns and not more
than about 12.5 microns; preferably, the thickness is about 6.0
microns. "Dry film" thickness refers to the thickness of a cured
film cast from a coating formulation onto the surface of the
membrane by standard coating techniques
Suitable hydrogel domains are formed of a curable copolymer of a
urethane polymer and a hydrophilic film-forming polymer.
Particularly preferred coatings are formed of a polyurethane
polymer having anionic carboxylate functional groups and non-ionic
hydrophilic polyether segments, which is crosslinked in the present
of polyvinylpyrrolidone and cured at a moderate temperature of
about 50.degree. C.
B. Electrolyte Phase
The electrolyte phase is a free-fluid phase including a solution
containing at least one compound, usually a soluble chloride salt,
that conducts electric current. The electrolyte phase flows over
the electrodes and is in contact with the hydrogel domain. The
devices of the present invention contemplate the use of any
suitable electrolyte solution, including standard, commercially
available solutions.
Generally speaking, the electrolyte phase should have the same or
less osmotic pressure than the sample being analyzed. In preferred
embodiments of the present invention, the electrolyte phase
includes normal saline.
C. Membrane Architectures
Prior art teaches that an enzyme containing membrane that resides
above an amperometric electrochemical sensor can possess the same
architecture throughout the electrode surfaces. However, the
function of converting glucose into hydrogen peroxide by glucose
oxidase may only by necessary above the working electrode. In fact,
it may be beneficial to limit the conversion of glucose into
hydrogen peroxide above the counter electrode. Therefore, the
present invention contemplates a number of membrane architectures
that include a multi-region membrane wherein the regions include at
least one domain.
Referring now to FIG. 2A, which shows one desired embodiment of the
general architecture of a three region membrane, first region 33 is
permeable to oxygen and glucose and includes a cell disruptive
domain distant from the electrodes and a cell impermeable domain
less distant from the electrodes. The second region 31 is permeable
to oxygen and includes a glucose exclusion domain and region three
32 includes a resistance domain, distant from the electrochemically
reactive surfaces, an immobilized enzyme domain less distant from
the electrochemically reactive surfaces, an interference domain
less distant from the electrochemically reactive surfaces than the
immobilized enzyme and a hydrogel domain adjacent to the
electrochemically reactive surfaces. The multi-region membrane is
positioned over the sensor interface 30 of the non-conductive body
10, covering the working electrode 21, the reference electrode 20
and the counter electrode 22. The electrodes are brazed to the
sensor head and back filled with epoxy 28.
In FIG. 2B, the glucose exclusion domain has been positioned over
the electrochemically reactive surfaces such that it does not cover
the working electrode 21. To illustrate this, a hole 35 has been
created in the second region 31 and positioned directly above the
working electrode 21. In this way, glucose is blocked from entering
the underlying enzyme membrane above the counter electrode 22 and
O.sub.2 is conserved above the counter electrode because it is not
being consumed by the glucose oxidation reaction. The
glucose-blocking domain is made of a material that allows
sufficient O.sub.2 to pass to the counter electrode. The
glucose-blocking domain may be made of a variety of materials such
as silicone or silicone containing copolymers. Preferably, the
glucose-blocking domain is made of silicone.
In FIG. 2C, the multi-region membrane is shown as being constructed
of two regions: a first region 33 which includes a cell disruptive
domain and a cell impermeable domain; and a further region 32.
Region 32 is defined herein as including an enzyme immobilized
domain, interference domain, and hydrogel domain and may also
include a resistance domain. Region 32 is referred to as the "third
region" in embodiments where the multi-region membrane includes
three regions. In the embodiment shown, a silicone domain plug 36
positioned over the counter electrode 22 in order to eliminate the
consumption of O.sub.2 above the counter electrode by the oxidation
of glucose with glucose oxidase. The enzyme immobilized domain can
be fabricated as previously described, then a hole punched into the
domain. The silicone domain plug 36 may be cut to fit the hole, and
then adhered into place, for example, with silicone adhesive (e.g.,
MED-1511, NuSil, Carpinteria, Calif.).
In FIG. 2D, the immobilized enzyme domain of the multi-region
membrane can be fabricated such that active enzyme 37 is positioned
only above the working electrode 21. In this architecture, the
immobilized enzyme domain may be prepared so that the glucose
oxidase only exists above the working electrode 21. During the
preparation of the multi-region membrane, the immobilized enzyme
domain coating solution can be applied as a circular region similar
to the diameter of the working electrode. This fabrication can be
accomplished in a variety of ways such as screen printing or pad
printing. Preferably, the enzyme domain is pad printed during the
enzyme membrane fabrication with equipment as available from Pad
Print Machinery of Vermont (Manchester, Vt.). These architectures
eliminate the consumption of O.sub.2 above the counter electrode 22
by the oxidation of glucose with glucose oxidase.
In FIG. 2E, the immobilized enzyme of the multi-region membrane in
region 32 may be deactivated 38 except for the area covering the
working electrode 21. In some of the previous membrane
architectures, the glucose oxidase is distributed homogeneously
throughout the immobilized enzyme domain. However, the active
enzyme need only reside above the working electrode. Therefore, the
enzyme may be deactivated 38 above the counter 22 and reference 20
electrodes by irradiation. A mask that covers the working electrode
21, such as those used for photolithography can be placed above the
membrane. In this way, exposure of the masked membrane to
ultraviolet light deactivates the glucose oxidase in all regions
except that covered by the mask.
FIG. 2F shows an architecture in which the third region 32 which
includes immobilized enzyme only resides over the working electrode
21. In this architecture, consumption of O.sub.2 above the counter
electrode 22 by the oxidation of glucose with glucose oxidase is
eliminated.
D. The Electrode Assembly
The electrode assembly of this invention comprises a non-conductive
body and three electrodes affixed within the body having
electrochemically reactive surfaces at one location on the body and
an electronic connection means at another location on the body and
may be used in the manner commonly employed in the making of
amperometric measurements. A sample of the fluid being analyzed is
placed in contact with a reference electrode, e.g.,
silver/silver-chloride, a working electrode which is preferably
formed of platinum, and a counter electrode which is preferably
formed of gold or platinum. The electrodes are connected to a
galvanometer or polarographic instrument and the current is read or
recorded upon application of the desired D.C. bias voltage between
the electrodes.
The ability of the present device electrode assembly to accurately
measure the concentration of substances such as glucose over a
broad range of concentrations in fluids including undiluted whole
blood samples enables the rapid and accurate determination of the
concentration of those substances. That information can be employed
in the study and control of metabolic disorders including
diabetes.
The present invention contemplates several structural architectures
that effectively increase the electrochemically reactive surface of
the counter electrode. In one embodiment, the diameter of wire used
to create the counter electrode is at least twice the diameter of
the working electrode. In this architecture, it is preferable that
the electrochemically reactive surface of the counter electrode be
not less than about 2 and not more than about 100 times the surface
area of the working electrode. More preferably, the
electrochemically reactive surface of the counter electrode is not
less than about 2 and not more than about 50, not less than about 2
and not more than about 25 or not less than about 2 and not more
than about 10 times the surface area of the working electrode. In
another embodiment, the electrochemically reactive surface is
larger that the wire connecting this surface to the sensor head. In
this architecture, the electrode could have a cross-sectional view
that resembles a "T". The present invention contemplates a variety
of configurations of the electrode head that would provide a large
reactive surface, while maintaining a relatively narrow connecting
wire. Such configurations could be prepared by micromachining with
techniques such as reactive ion etching, wet chemical etching and
focused ion beam machining as available from Norsam Technologies
(Santa Fe, N.Mex.).
In another embodiment, the diameter of the counter electrode is
substantially similar to the working electrode; however, the
surface of the counter electrode has been modified to increase the
surface area such that it has at least twice the surface area of
the working electrode. More specifically the counter electrodes
surface may be textured, effectively increasing its surface area
without significantly increasing its diameter. This may be
accomplished by a variety of methods known to those skilled in the
art including, such as acid etching. The electrochemically reactive
surface may be provided in a variety of shapes and sizes (e.g.
round, triangular, square or free form) provided that it is at
least twice the surface area of the working electrode.
In all of the architectures described, the electrodes are prepared
from a 0.020'' diameter wire having the desired modified reactive
surface. The electrodes are secured inside the non-conductive body
by brazing. The counter electrode is preferably made of gold or
platinum.
III. Analyte Measuring Device
A preferred embodiment of an analyte measuring device including a
sensor head according to the present invention is shown in FIG. 6A.
In this embodiment, a ceramic body 1 and ceramic head 10 houses the
sensor electronics that include a circuit board 2, a microprocessor
3, a battery 4, and an antenna 5. Furthermore, the ceramic body 1
and head 10 possess a matching taper joint 6 that is sealed with
epoxy. The electrodes are subsequently connected to the circuit
board via a socket 8.
As indicated in detail in FIG. 6B, three electrodes protrude
through the ceramic head 10, a platinum working electrode 21, a
platinum counter electrode 22 and a silver/silver chloride
reference electrode 20. Each of these is hermetically brazed 26 to
the ceramic head 10 and further secured with epoxy 28. The sensing
region 24 is covered with a multi-region membrane described above
and the ceramic head 10 contains a groove 29 so that the membrane
may be affixed into place with an o-ring.
IV. Experimental
The following examples serve to illustrate certain preferred
embodiments and aspects of the present invention and are not to be
construed as limiting the scope thereof
In the preceding description and the experimental disclosure which
follows, the following abbreviations apply: Eq and Eqs
(equivalents); mEq (milliequivalents); M (molar); mM (millimolar)
.mu.M (micromolar); N (Normal); mol (moles); mmol (millimoles);
.mu.mol (micromoles); nmol (nanomoles); g (grams); mg (milligrams);
.mu.g (micrograms); Kg (kilograms); L (liters); mL (milliliters);
dL (deciliters); .mu.L (microliters); cm (centimeters); mm
(millimeters); .mu.m (micrometers); nm (nanometers); h and hr
(hours); min. (minutes); s and sec. (seconds); .degree. C. (degrees
Centigrade); Astor Wax (Titusville, Pa.); BASF Wyandotte
Corporation (Parsippany, N.J.); Data Sciences, Inc. (St. Paul,
Minn.); DuPont (DuPont Co., Wilmington, Del.); Exxon Chemical
(Houston, Tex.); GAF Corporation (New York, N.Y.); Markwell Medical
(Racine, Wis.); Meadox Medical, Inc. (Oakland, N.J.); Mobay (Mobay
Corporation, Pittsburgh, Pa.); NuSil Technologies (Carpenteria,
Calif.) Sandoz (East Hanover, N.J.); and Union Carbide (Union
Carbide Corporation; Chicago, Ill.).
EXAMPLE 1
Preparation of the Multi-region Membrane
A. Preparation of the First Region
The cell disruptive domain may be an ePTFE filtration membrane and
the cell impermeable domain may then be coated on this domain
layer. The cell impermeable domain was prepared by placing
approximately 706 gm of dimethylacetamide (DMAC) into a 3 L
stainless steel bowl to which a polycarbonateurethane solution
(1325 g, Chronoflex AR 25% solids in DMAC and 5100 cp) and
polyvinylpyrrolidone (125 g, Plasdone K-90 D) are added. The bowl
was then fitted to a planetary mixer with a paddle type blade and
the contents were stirred for 1 hour at room temperature. This
solution was then coated on the cell disruptive domain by knife
edge drawn at a gap of 0.006'' and dried at 60.degree. C. for 24
hours.
Alternatively, the polyurethane polyvinylpyrrolidone solution
prepared above can be coated onto a PET release liner using a knife
over roll coating machine. This material is then dried at
305.degree. F. for approximately 2 minutes. Next the ePTFE membrane
is immersed in 50:50 (w/v) mixture of THF/DMAC and then placed atop
the coated polyurethane polyvinylpyrrolidone material. Light
pressure atop the assembly intimately embeds the ePTFE into the
polyurethane polyvinylpyrrolidone. The membrane is then dried at
60.degree. C. for 24 hours.
B. Preparation of the Glucose Exclusion Domain
An oxime cured silicone dispersion (NuSil Technologies, MED-6607)
was cast onto a polypropylene sheet and cured at 40.degree. C. for
three days.
C. Preparation of the Third Region
The "third region" or "further region" includes a resistance
domain, an immobilized enzyme domain, an interference domain and an
hydrogel domain. The resistance domain was prepared by placing
approximately 281 gm of dimethylacetamide into a 3 L stainless
steel bowl to which a solution of polyetherurethaneurea (344 gm of
Chronothane H, 29,750 cp at 25% solids in DMAC). To this mixture
was added another polyetherurethaneurea (312 gm, Chronothane 1020,
6275 cp at 25% solids in DMAC.) The bowl was fitted to a planetary
mixer with a paddle type blade and the contents were stirred for 30
minutes at room temperature. The resistance domain coating
solutions produced is coated onto a PET release liner (Douglas
Hansen Co., Inc. Minneapolis, Minn.) using a knife over roll set at
a 0.012'' gap. This film is then dried at 305.degree. F. The final
film is approximately 0.0015'' thick.
The immobilized enzyme domain was prepared by placing 304 gm
polyurethane latex (Bayhydrol 140 AQ, Bayer, Pittsburgh, Pa.) into
a 3 L stainless steel bowl to which 51 gm of pyrogen free water and
5.85 gm of glucose oxidase (Sigma type VII from Aspergillus niger)
is added. The bowl was then fitted to a planetary mixer with a
whisk type blade and the mixture was stirred for 15 minutes.
Approximately 24 hr prior to coating a solution of glutaraldehyde
(15.4 mL of a 2.5% solution in pyrogen free water) and 14 mL of
pyrogen free water was added to the mixture. The solution was mixed
by inverting a capped glass bottle by hand for about 3 minutes at
room temperature. This mixture was then coated over the resistance
domain with a #10 Mayer rod and dried above room temperature
preferably at about 50.degree. C.
The interference domain was prepared by placing 187 gm of
tetrahydrofuran into a 500 mL glass bottle to which an 18.7 gm
aliphatic polyetherurethane (Tecoflex SG-85A, Thermedics Inc.,
Woburn, Mass.) was added. The bottle was placed onto a roller at
approximately 3 rpm within an oven set at 37.degree. C. The mixture
was allowed to roll for 24 hr. This mixture was coated over the
dried enzyme domain using a flexible knife and dried above room
temperature preferably at about 50.degree. C.
The hydrogel domain was prepared by placing 388 gm of polyurethane
latex (Bayhydrol 123, Bayer, Pittsburgh, Pa. in a 3 L stainless
steel bowl to which 125 gm of pyrogen free water and 12.5 gm
polyvinylpyrrolidone (Plasdone K-90D) was added. The bowl was then
fitted to a planetary mixer with a paddle type blade and stirred
for 1 hr at room temperature. Within 30 minutes of coating
approximately 13.1 mL of carbodiimide (UCARLNK) was added and the
solution was mixed by inverting a capped polyethylene jar by hand
for about 3 min at room temperature. This mixture was coated over
the dried interference domain with a #10 Mayer rod and dried above
room temperature preferably at about 50.degree. C.
In order to affix this multi-region membrane to a sensor head, it
is first placed into buffer for about 2 minutes. It is then
stretched over the nonconductive body of sensor head and affixed
into place with an o-ring.
EXAMPLE 2
In vitro Evaluation of Sensor Devices
This example describes experiments directed at sensor function of
several sensor devices contemplated by the present invention.
In vitro testing of the sensor devices was accomplished in a manner
similar to that previously described. [Gilligan et al., Diabetes
Care 17:882-887 (1994)]. Briefly, devices were powered on and
placed into a polyethylene container containing phosphate buffer
(450 ml, pH 7.30) at 37.degree. C. The container was placed onto a
shaker (Lab Line Rotator, model 1314) set to speed 2. The sensors
were allowed to equilibrate for at least 30 minutes and their
output value recorded. After this time, a glucose solution (9.2 ml
of 100 mg/ml glucose in buffer) was added in order to raise the
glucose concentration to 200 mg/dl within the container. The
sensors were allowed to equilibrate for at least 30 minutes and
their output value recorded. Again, a glucose solution (9.4 ml of
100 mg/ml glucose in buffer) was added in order to raise the
glucose concentration to 400 mg/dl within the container. The
sensors were allowed to equilibrate for at least 30 minutes and
their output value recorded. In this way, the sensitivity of the
sensor to glucose is given as the slope of sensor output versus
glucose concentration. The container was then fitted with an
O.sub.2 meter (WTW, model Oxi-340) and a gas purge. A mixture of
compressed air and nitrogen was used to decrease the O.sub.2
concentration. Sensor output was recorded at an ambient O.sub.2
level, then sensor output was recorded for the following O.sub.2
concentrations; 1 mg/L, 0.85 to 0.75 mg/L, 0.65 to 0.55 mg/L and
0.40 to 0.30 mg/L. In this way, the function of the sensor could be
compared to its function at ambient O.sub.2 .
Sensor devices like the one shown in FIGS. 6A and 6B, which
included inventive sensor heads having a multi-region membrane with
the architecture shown in FIG. 2B, were tested in vitro. Eight of
these devices were fitted with membranes that possessed a 0.020''
diameter hole, four with a 0.0015'' thick polyurethane (Chronoflex
AR, CardioTech International Inc.) and four with a 0.032'' thick
silicone (MED-1511, NuSil Technologies Inc.). The hole was
positioned above the working electrode and both membranes were
secured to the device with an o-ring. Four control devices were
also tested which were fitted with a multi-region membrane which
lacked region 31 shown in FIB. 2B.
As discussed above, for oxygen to be consumed in the sensing region
32 above the electrodes, glucose is required. By placing region 31
shown in FIG. 2B, which includes a glucose blocking domain, above
all areas other than above the working electrode 21, oxygen
consumption in areas other than working electrode areas is limited.
In contrast, by eliminating region 31 in the control devices, less
overall oxygen becomes available to electrode surfaces due to the
increased availability of glucose.
The devices were activated, placed into a 500 ml-polyethylene
container with sodium phosphate buffered solution (300 ml, pH 7.3)
and allowed to equilibrate. Each device's baseline value was
recorded. Then 12 ml of glucose solution (100 mg/ml in sodium
phosphate buffer) was added to the container so that the total
glucose concentration became 400 mg/dL. After this, the container
was covered and fitted with an oxygen sensor and a source of
nitrogen and compressed air. In this way, the oxygen concentration
was controlled with a gas sparge. A glucose value was recorded for
each device at decreasing oxygen concentrations from ambient to
approximately 0.1 mg/L.
FIG. 7 graphically represents the formation of a device of the
present invention utilizing the multi-region membrane architecture
in FIG. 2B in vitro. The data is expressed in percent Device
Function at 400 mg/dL glucose vs. oxygen concentration. The percent
function of the device is simply the device output at any given
oxygen concentration divided by that device's output at ambient
oxygen. The results from FIG. 7 indicate that inventive sensor
devices containing the silicone membrane have better function at
lower oxygen concentrations relative to both the control devices
and the devices containing the polyurethane membrane. For example,
at an oxygen concentration of about 0.5 mg/L, devices containing
the silicone membrane are providing 100% output as compared to 80%
output for the control devices.
EXAMPLE 3
The Effect of Varying the Size and Material of the Counter
Electrode on Sensor Response and Accuracy
An in vitro testing procedure used in this example was similar to
that described in Example 2. Six devices similar to the one shown
in FIGS. 6A and 6B were fitted with the multi-region membrane
described herein. Two of these tested devices were comparative
devices that possessed Pt counter electrodes having a 0.020''
diameter; this diameter provided for an electrochemically reactive
surface of the counter electrode which was substantially equal to
the surface area of the working electrode, as schematically shown
in FIG. 8. In FIG. 8, the electrode-membrane region includes two
distinct regions, the compositions and functions of which have
already been described. Region 32 includes an immobilized enzyme.
Region 33 includes a cell disruptive domain and a cell impermeable
domain. The top ends of electrodes 21 (working), 20 (reference) and
22 (counter) are in contact with an electrolyte phase 30, a
free-flowing phase. Two other tested devices possessed Pt counter
electrodes having a 0.060'' diameter. Finally, two additional
devices possessed Au counter electrodes having a 0.060'' diameter.
The 0.006'' diameter devices provided for an electrochemically
reactive surface of the counter electrode which was approximately
six times the surface area of the working electrode. Each of the
devices including counter electrodes of 0.060'' diameter include a
multi-region membrane above the electrode region which is similar
to that shown in FIG. 8.
The devices were activated, placed into a 500 ml-polyethylene
container with sodium phosphate buffered solution (300 ml, pH 7.3)
and allowed to equilibrate. Each device's baseline value was
recorded. Then 12 ml of glucose solution (100 mg/ml in sodium
phosphate buffer) was added to the container so that the total
glucose concentration became 400 mg/dL. After this, the container
was covered and fitted with an oxygen sensor and a source of
nitrogen and compressed air. In this way, the oxygen concentration
was controlled with a gas sparge. A counter electrode voltage was
recorded for each device at decreasing oxygen concentrations from
ambient to approximately 0.1 mg/L.
FIG. 9 graphically presents the counter electrode voltage as a
function of oxygen concentration and 400 mg/dL glucose. This figure
demonstrates that both the large Pt and Au counter electrode
devices do not begin to reach the circuitry limits at low oxygen
concentrations. Therefore, increased performance and accuracy can
be obtained from a counter electrode that has an electrochemical
reactive surface greater than the surface area of the working
electrode.
The description and experimental materials presented above are
intended to be illustrative of the present invention while not
limiting the scope thereof It will be apparent to those skilled in
the art that variations and modifications can be made without
departing from the spirit and scope of the present invention.
* * * * *